Delivery systems based on hydrogel compositions and methods thereof

ABSTRACT

The invention provides a novel, versatile degradable hydrogel composition, and methods thereof, with precisely tunable stiffness, plasticity (e.g., degree of covalent vs. physical crosslinks) and predictive disintegration rates degradation, allowing controlled disintegration and release of therapeutic cells or pharmaceuticals and/or in vitro 3D cell expansion.

PRIORITY CLAIMS AND RELATED APPLICATIONS

This application claims the benefit to U.S. Provisional Application No.62/584,402, filed Nov. 10, 2017, the entire content of which isincorporated herein by reference for all purposes.

TECHNICAL FIELD OF THE INVENTION

The invention generally relates to hydrogels and delivery oftherapeutics. More particularly, the invention relates to a novel,versatile degradable hydrogel composition, and methods thereof, havingprecisely tunable stiffness, plasticity and degradation, allowingcontrolled disintegration and release of therapeutic cells orpharmaceuticals.

BACKGROUND OF THE INVENTION

Hydrogels have been used to encapsulate and deliver stem cells orprimary cells to promote tissue regeneration. Hydrogel with controlleddegradation behaviors are especially useful for a variety of biomedicalapplications (e.g., drug delivery and tissue regeneration). It isincreasingly appreciated that hydrogel stiffness, matrixplasticity/viscoelasticity and degradation modulate cell-hydrogelinteractions and consequently cellular behaviors includingproliferation, morphogenesis, differentiation/phenotypical matrixdeposition. (Diekjurgen, et al. 2017 Biomaterials 141, 96; Shin, et al.2016 Proc. Natl. Acad. Sci. U.S.A 113 (43), 12126; Yang, et al. 2014Nature materials 13 (6), 645; Khetan, et al. 2013 Nature materials 12(5), 458.)

Natural polymer-based hydrogels such as Matrigel™ and a range ofpolysaccharide-based hydrogels such as alginate have been exploited forcell encapsulations due to their cytocompatibility. Recent studiesdemonstrated the impact of viscoelasticity of alginate-based hydrogelson the cellular behavior of encapsulated mesenchymal stem cells (MSCs)by modulating the molecular weight of alginate, the degree ofCa²⁺-induced physical crosslinking, and the introduction of covalentlytethered poly(ethylene glycol) chains. Viscoelasticity of naturalhydrogels, imposed by extensive physical interactions within the polymernetwork, has been recognized as critical for alleviating the elasticstress of hydrogel matrix to encapsulated cells. (Diekjurgen, et al.2017 Biomaterials 141, 96; Chaudhuri, et al. 2016 Nature materials 15(3), 326.)

Difficulty in achieving regiospecific and stoichiometrically controlledchemical modification of natural polymers limits tunability of theirbiophysical and degradative properties. This limitation, combined withbatch-to-batch variation in their matrix compositions, residueanimal-derived components, and risks for contamination/immunogenicity,presents significant hurdles to their regulatory approval and clinicaltranslation.

Unlike natural biopolymer-based hydrogels, wholly synthetic hydrogelscan be prepared free of biocontaminants. The use of wholly synthetichydrogels for cell encapsulation, however, has been largely limited toinvestigating the impact of stiffness of photo-crosslinkedpolymethacrylate-based hydrogels on the cellular fate of MSCs or thematrix deposition of encapsulated chondrocytes in vitro. Modulation ofstiffness of these hydrogels was mainly accomplished by altering degreeof covalent crosslinking or polymer weight fractions, which did notaddress the negative impact of high elastic stress imposed by thesehydrogel networks to the metabolism of encapsulated cells. (Benoit, etal. 2008 Nature materials 7 (10), 816; Lutolf, et al. 2005 Nat.Biotechnol. 23 (1), 47; Mao, et al. 2016 Biomaterials 98, 184; Engler,et al. 2006 Cell 126 (4), 677; Butler, et al. 2009 Tissue engineering.Part B, Reviews 15 (4), 477; Bryant, et al. 2004 J. Orthop. Res. 22 (5),1143.)

In addition, the radical initiators and photo-irradiation employed tocovalently crosslink these hydrogels are known for imposingcytotoxicity/altering gene expression of encapsulated cells. Meanwhile,covalent incorporation of degradable polylactide segments has beenutilized to introduce degradability to elastic covalently crosslinkedhydrogel network to promote encapsulated cell matrix deposition and/orto enable cell release. Such a method of modulating hydrogeldegradability, however, is empirical rather than predictive in nature,and generates significant inflammatory acidic degradation products.(Gasparian, et al. 2015 Anal. Biochem. 484, 1; Fedorovich, et al. 2009Biomaterials 30 (3), 344; Filion, et al. 2011 Biomaterials 32 (4), 985;Bergsma 1995 Biomaterials 16 (1), 25; Chu, et al. 2017 Tissue Eng Part A23 (15-16), 795; Bryant, et al. 2003 J. Biomed. Mater. Res. A 64 (1),70.)

Overall, it remains a significant challenge to develop a whollysynthetic 3D hydrogel where its matrix stiffness, plasticity anddegradative property can be quantitatively and predictively tuned over abroad range for facile cell encapsulation, to accommodate cellproliferation/phenotypical matrix deposition, and to enable preciselytimed release for in vivo cell delivery.

There is an urgent and ongoing need for novel and improved approachesthat effectively address these issues.

SUMMARY OF THE INVENTION

The invention provides wholly synthetic hydrogels useful for cellencapsulation and delivery with predictive tuning of stiffness,plasticity/viscoelasticity and degradation of tissue matrices regulatecell behavior. In particular, disclosed is a novel 3D synthetic hydrogelplatform with explicitly controlled ratio of biorthogonal covalent vs.non-covalent crosslinking of cytocompatible building blocks andstrategic placement of a single stable vs. labile linker near thecrosslinking site. The former dictates the matrix stiffness andviscoelasticity while the later predicts matrix degradation.

For example, hydrogels with varying stiffness (e.g., 0.86-11.75 KPa) andmatrix plasticity (degree of covalent vs. physical crosslinks) andpredictive disintegration rates (e.g., 18->150 days) were prepared from2 pairs of labile/stable building blocks in varying ratios for 3Dencapsulation of rodent and human chondrocytes. Stiffer hydrogelsstrengthened by dynamic physical crosslinks between dibenzocyclooctyne(DBCO)-terminated building blocks better absorbed the stress andaccommodated the expanding volume imposed by the proliferation of andchondrogenic matrix deposition by encapsulated chondrocytes. Degradationof the hydrogel promoted the proliferation and matrix deposition ofencapsulated chondrocytes, with those released upon timed geldisintegration maintaining their chondrogenic phenotype.

The present disclosure demonstrates that high-fidelity biorthogonalcovalent and physical crosslinking of a small set of designer buildingblocks can be conveniently exploited to engineer 3D synthetic nicheswith tunable matrix stiffness, plasticity and predictive degradativeproperties for cell encapsulation, expansion and release

In one aspect, the invention generally relates to a hydrogel comprisinga 3-dimentional crosslinking network of a combination of covalentcrosslinking (CX) and non-covalent crosslinking (NCX) of hydrophilicbranched polymers,

wherein the hydrophilic branched polymers are star-branched with atleast 3 arms.

In another aspect, the invention generally relates to a hydrogelcomposition comprising a hydrogel disclosed herein and a biologicallyactive payload encapsulated therein.

In yet another aspect, the invention generally relates to a device orimplant that includes a hydrogel or a hydrogel composition disclosedherein.

In yet another aspect, the invention generally relates to a method formodulating one or more properties of a hydrogel. The method includes:incorporating in the hydrogel a 3-dimentional crosslinking network of acombination of covalent crosslinking (CX) and non-covalent crosslinking(NCX) of hydrophilic branched polymers,

adjusting and/or controlling the ratio of covalent crosslinking (CX) tonon-covalent crosslinking (NCX) of hydrophilic branched polymers; andadjusting and/or controlling the placement and ratio of one or morelabile linkages to one or more stable linkages in the hydrophilicbranched polymers. The hydrophilic branched polymers are star-branchedwith at least 3 arms. The property is one or more selected fromviscoelasticity, stiffness and degradation.

In yet another aspect, the invention generally relates to a method fordelivering a biologically active payload. The method includes: placingin a subject in need thereof a device or implant comprising a hydrogelor hydrogel composition disclosed herein, and causing a controlledrelease of the biologically active payload.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1. Depiction of catalyst-free, irradiation-free encapsulation ofcells by ClickGels with varying degrees of SPAAC crosslinks and physicalcrosslinks formed by mixing non-degradable or degradable azide- andDBCO-terminated 4-armPEG macromer building blocks in various ratios.

FIG. 2. Non-degradable ClickGels (5 w/v %) crosslinked betweenmismatched ratios of DBCO- and N₃-terminated macromers exhibit tunableswelling behavior, stiffness and viscoelasticity, with stiffer ClickGelsformed with excess DBCO-terminated macromers exhibiting faster stressrelaxation. (a) Swelling ratio. (b) Compressive moduli (0-30% strain) ofClickGels crosslinked between varying ratios of DBCO- and N₃-terminatedmacromers. (c) Compressive modulus changes (0-30% and 60-65% stainranges) upon the addition and removal of polyaromatic dye (bromophenolblue sodium salt). (d) Representative stress relaxation profiles ofClickGels composed of varying ratios of DBCO- and N₃-terminatedmacromers. (e) Stress relaxation time (τ_(1/2)) of varying ratios ofDBCO- and N₃-terminated macromers. ns: p>0.05; *p<0.05; **p<0.01;***p<0.001; *****p<0.0001 (a, b: one-way ANOVA with Dunnett's multiplecomparisons vs. the 1:1 formulation; c & e: two-way and one-way ANOVAwith Tukey's multiple comparisons, respectively). (f) Shear stressrelaxation time of a ClickGel with [amide-DBCO]:[CH2-N₃]=1:0.6determined as 1268.2±304.7 (mean±S.D.). ClickGel disc (8-mm in diameter,n=5) was equilibrated in PBS for 24 hours. Strain sweep and frequencysweep test were performed on an AG2000 rheometer to determine the linearviscoelastic region. Shear stress relaxation time was measured upon theapplication of a constant shear strain of 10% (5 s rise time).

FIG. 3. Stiffer non-degradable ClickGels (5 w/v %) with dynamicDBCO-DBCO physical crosslinks better accommodate the proliferation andchondrogenic ECM deposition of encapsulated iMACs over time. (a)Depiction of reversible formation of DBCO-DBCO crosslinks helpingdissipate stress imposed by the expanding cell mass ECM deposition byencapsulated cells. (b) Viability of encapsulated iMACs over time. (c)Temporal changes in compressive moduli (0-30% strain) of iMAC-ladenClickGels over 28-day culture. (d) Live (green)/dead (red) staining andtype II collagen (green)/DAPI (blue) immunofluorescent staining ofiMAC-laden ClickGels after 28-day culture. 250,000 iMACs wereencapsulated in 25-μL 5 wt % non-degradable ClickGels of varyingmacromer ratios and cultured in expansion media. ns: p>0.05; **p<0.01(two-way ANOVA with Sidak's multiple comparisons).

FIG. 4. Polymer contents further modulate stiffness of non-degradableClickGels crosslinked between varying ratios of DBCO- and N3-terminatedmacromers and the cellular proliferation of encapsulated iMACs overtime. a. Compressive moduli (0-30% strain) as a function of ClickGelpolymer content and ratios of DBCO- and N3-terminated macromers. b-d.Viability of encapsulated iMACs over time as a function of ClickGelpolymer content in ClickGels crosslinked from different ratios (b.0.6:1; c. 1:1; d. 1:0.6) of DBCO- and N₃-terminated macromers. 250,000iMACs were encapsulated in 25-μL 2.5, 5 or 10 wt % non-degradableClickGels of varying macromer ratios and cultured in expansion media.ns: p>0.05; *p<0.05; **p<0.01; ***p<0.001; *****p<0.0001 (two-way ANOVAwith Tukey's multiple comparisons vs. the 5 w/v % gel).

FIG. 5. Human chondrocytes encapsulated in non-degradable ClickGelsmaintain long-term viability and chondrocyte phenotype regardless of theratio of DBCO- and N3-terminated macromers. a. Viability of encapsulatedhuman chondrocytes as a function of ClickGel composition over 56 days.b. type II collagen (green)/DAPI (blue), type X collagen (green)/DAPI(blue) immunofluorescent staining and toluidine blue (for GAG) stainingof human chondrocyte-laden ClickGels of varying compositions on day 56of chondrogenic culture. 500,000 human chondrocytes were encapsulated in25-μL 5 wt % non-degradable ClickGels of varying macromer ratios andcultured in chondrogenic media (high glucose DMEM, 40-μg/mL L-proline,100-μg/mL sodium pyruvate, 1% insulin-transferrin-selenous acid mixture,100-nM dexamethasone and 10-ng/mL TGF-β3). ns: p>0.05 (two-way ANOVAwith Tukey's multiple comparisons vs. the 1:1 formulation at a giventime).

FIG. 6. ClickGel degradation enhances proliferation and chondrogenic ECMdepositions of encapsulated iMACs and human chondrocytes. a.Disintegration time of ClickGels (5 w/v %) formed between non-labileDBCO-terminated macromer and a mixture of non-labile and labileazide-terminated macromers upon incubation in expansion media. b.Viability of iMACs encapsulated within perfectly SPAAC-crosslinkedClickGels with varying degradability (250,000/25-pt gel) in expansionmedia over time. c. Live (green)/dead (red) staining and type IIcollagen (green)/DAPI (blue) immunofluorescent staining of iMAC-ladenClickGels with varying degradability on day 28 of culture in expansionmedia. d. Type II collagen (green)/DAPI (blue), type X collagen(green)/DAPI (blue) immunofluorescent staining and toluidine blue (forGAG) staining of human chondrocyte-laden ClickGels with varyingdegradability in chondrogenic media. ns: p>0.05; *p<0.05; **p<0.01;*****p<0.0001 (two-way ANOVA with Tukey's multiple comparisons vs. the100% gel).

FIG. 7. Degradable ClickGels (5 w/v %, perfectly SPAAC-crosslinked)enable timed release of encapsulated iMACs with retained chondrogenicphenotype. a. Depiction of ClickGel network disintegration as a resultof hydrolysis of labile SPAAC crosslinks and the release of encapsulatedcells. b. Crystal violet staining of cells released from ClickGels withvarying degradability over different culture duration in expansionmedia. c. GAG and type II collagen staining of the pellet of releasediMACs. Upon complete disintegration of the degradable ClickGel, releasediMAC were pelleted and cultured in expansion media for 10 days prior tostaining.

FIG. 8: Non-degradable ClickGels (5 w/v %) crosslinked betweenmismatched ratios of DBCO- and N₃-terminated macromers exhibit varyingstiffness, with the ClickGel formed with only 30% covalent SPAACcrosslinks and 70% excess DBCO-terminated macromers exhibiting highercompressive moduli (0-30% strain), viability than matrix deposition thanthose formed with 60-80% covalent SPAAC crosslinks but excessN₃-terminated macromers. a. compressive moduli. b. Viability ofencapsulated iMAC as a function of ClickGel composition over 28 days. c.Live (green)/dead (red) staining and type II collagen (green)/DAPI(blue) immunofluorescent staining of iMAC-laden ClickGels (0.8:1 vs1:0.3) on day 28 of culture in expansion media. ns: p>0.05; *p<0.05;***p<0.001 (one-way ANOVA with Dunnett's multiple comparisons or Sidak'smultiple comparisons test vs. the 1:0.3 formulation).

FIG. 9: iMACs proliferate and maintain chondrocyte phenotype inperfectly SPAAC-crosslinked non-degradable ClickGel while excessiveinitial cell encapsulation densities reduce long-term viability ofiMACs. a. Viability of iMACs over time as a function of initialencapsulation density. b. Viability of iMACs on day 56 as a function ofinitial encapsulation density. c. Live (green)/dead (red) staining, typeII collagen (green)/DAPI (blue) staining, type X collagen (green)/DAPI(blue) immunofluorescent staining and toluidine blue staining ofiMAC-laden ClickGels with varying initial cell encapsulation densitiesafter 56-day culture. 250,000-1,000,000 iMACs were encapsulated in 25-μL5 wt % perfectly SPAAC-crosslinked non-degradable ClickGels and culturedin expansion media (high glucose DMEM supplemented with 10% FBS and 1%Pen/Strep). ***p<0.001; *****p<0.0001 (one-way ANOVA with Tukey'smultiple comparisons vs. the 250K cell encapsulation density).

FIG. 10: ClickGel degradation enhances proliferation and chondrogenicECM depositions of encapsulated iMACs in imperfect network a. Viabilityof iMACs encapsulated within perfectly SPAAC-crosslinked ClickGels withvarying degradability (250,000/25 μL gel) in expansion media over time.b. Live (green)/dead (red) staining, c. type II collagen (green)/DAPI(blue) and d. type X collagen (green)/DAPI (blue) immunofluorescentstaining of iMAC-laden ClickGels with varying degradability on day 28 ofculture in expansion media. e. toluidine blue (for GAG) staining ofhuman chondrocyte-laden ClickGels with varying degradability inchondrogenic media. ns: p>0.05; *p<0.05; **p<0.01; *****p<0.0001(two-way ANOVA with Tukey's multiple comparisons vs. the 100% gel).

FIG. 11: Rat BMSC encapsulated in non-degradable ClickGels maintainviability and chondrocyte phenotype in chondrogenic culture. a.Viability of encapsulated rat BMSC as a function of ClickGel compositionover 28 days. b. type II collagen (green)/DAPI (blue) and type Xcollagen (green)/DAPI (blue) immunofluorescent staining of ratBMSC-laden ClickGels of varying compositions on day 28 of chondrogenicculture. 250,000 rat BMSCs were encapsulated in 25-μL 5 wt %non-degradable ClickGels of varying macromer ratios and cultured inchondrogenic media.

DETAILED DESCRIPTION OF THE INVENTION

The invention provides a novel 3D synthetic hydrogel platform withexplicitly controlled ratio of biorthogonal covalent vs. non-covalentcrosslinking of cytocompatible building blocks and strategic placementof a single stable vs. labile linker near the crosslinking site. Thewholly synthetic hydrogels disclosed herein are useful for cellencapsulation/delivery with predictive tuning of stiffness,plasticity/viscoelasticity and degradation of tissue matrices regulatecell behavior. In particular, covalent vs. non-covalent crosslinkingcytocompatible building blocks regulators matrix stiffness andviscoelasticity while positioning of stable vs. labile linkers controlsmatrix degradation.

The disclosed synthetic hydrogel is characterized by a well-structuredcrosslinking network where the matrix stiffness and plasticity isdictated by tunable combinations of biorthogonal covalent andnon-covalent physical crosslinking of cytocompatible building blockswhile the degradation is predictively controlled by the strategicplacement of isolated labile linkages near the crosslinking site. Forexample, strain-promoted alkyne-azide cycloaddition (SPAAC) can proceedefficiently in the absence of catalysts or irradiation underphysiological conditions between reactants functionalized with azidesand cyclooctynes, which are absent from native cellular/tissueenvironment. It has been utilized for the preparation of covalentlycrosslinked hydrogels for cell encapsulation. (Jewett, et al. 2010 Chem.Soc. Rev. 39 (4), 1272; Agard, et al. 2005 J. Am. Chem. Soc. 127 (31),11196; Zheng, et al. 2012 ACS macro letters 1 (8), 1071; Caldwell, etal. 2017 Advanced healthcare materials 6 (15); Xu, et al. 2014 J. Am.Chem. Soc. 136 (11), 4105; Xu, et al. 2011 Chem. Asian J. 6 (10), 2730;DeForest, et al. 2009 Nature materials 8 (8), 659.)

It was recently demonstrated that by SPAAC-crosslinking 4-armedpoly(ethylene glycol)-tetra-dibenzocyclooctyl (4-armPEG-DBCO) and4-armed poly(ethylene glycol)-tetra-azide (4-armPEG-azide) with a singlelabile ester linkage (X═O; Y═OC(O)—C₃H₆) or stable (X═NH; Y=absent)linkage near the azide or DBCO group (FIG. 1, [DBCO]: [N₃]=1:1),cytocompatible hydrogels named ClickGels with gelling kinetics (2-5 min)suitable for cell encapsulation and predictive degradation over a broadrange could be prepared. The strategic placement of a single labileester linkage at either side of the SPAAC crosslink (distinct K_(d)'sfor hydrolysis at X vs. Y site) within a well-structured hydrophilicnetwork enabled prediction of the ClickGel disintegration based on1^(st)-order hydrolytic cleavage kinetics. (Xu, et al. 2014 J. Am. Chem.Soc. 136 (11), 4105; Xu, et al. 2014 J. Am. Chem. Soc. 136 (11), 4105.)

While prospective tuning of the ClickGel disintegration rate from daysto months, was accomplished, the challenge remained for controlledengineering of stiffness and matrix plasticity/viscoelasticity.

The present invention solves the problem by quantitatively altering thedegrees of covalent SPAAC crosslinks vs. non-covalent crosslinks, forexample, driven by hydrophobic and H-bonding interactions betweenuntethered DBCO end groups (FIG. 1, bottom). Furthermore, the preciselytunable matrix stiffness, plasticity and degradative property ofClickGels are demonstrated to translate into an improved 3D synthetichydrogel allowing for long-term encapsulation and timed releaseproperties. This is of profound significance because matrix assistedchondrocyte delivery is a vital clinical treatment option for articularcartilage lesion, which is known for limited self-repair andregenerative capability due to its avascular, aneural nature and a denseECM that impedes autologous cell migration.

Cartilage is an anisotropic tissue with complex dynamic mechanicalproperties that change throughout development. The local stiffness ofthe ECM surrounding the chondrocytes is much lower than that of the bulkadult cartilage tissue. (Alexopoulos, et al. 2005 J. Biomech. 38 (3),509; Guilak, et al. 1999 Osteoarthritis Cartilage 7 (1), 59.) Previouswork suggests that low stiffness hydrogel scaffolds could help promotechondrocyte proliferation and maintain chondrogenic phenotype. (Wang, etal. 2017 Biomaterials 120, 11; Callahan, et al. 2013 Acta Biomater. 9(4), 6095.) However, minimal stiffness requirement of 3D hydrogelmatrices and the benefit of controlled introduction of matrixplasticity/viscoelasticity for ensuring viability, proliferation,chondrogenic ECM deposition of encapsulated chondrocytes has not beenestablished. The ability to weaken or strengthen ClickGel via precisetuning of the degrees of covalent SPAAC and DBCO-DBCO physicalcrosslinking, which may be dynamically disrupted and reformed toaccommodate cellular mass expansion, by mismatching the ratio of DBCO-vs. N₃-terminated macromers at a given overall polymer content offers anexciting opportunity to interrogate the impact of biomechanical cues onthe cellular behavior of encapsulated chondrocytes.

In one aspect, the invention generally relates to a hydrogel comprisinga 3-dimentional crosslinking network of a combination of covalentcrosslinking (CX) and non-covalent crosslinking (NCX) of hydrophilicbranched polymers,

wherein the hydrophilic branched polymers are star-branched with atleast 3 arms.

In certain embodiments, the physical interactions are selected from thegroup consisting of hydrophobic-hydrophobic interaction, π-π stacking,hydrogen bonding, electrostatic interaction, polar interaction, or acombination of one or more thereof.

In certain embodiments, the hydrophilic branched polymers are linked toNCX or CX via one or more labile linkages and/or one or more stablelinkages.

In certain embodiments, the one or more labile linkages are susceptibleto hydrolysis.

Exemplary labile linkages that are susceptible to hydrolysis includeesters, carbonates, orthoester, anhydride, and thioester.

In certain embodiments, the one or more labile linkages include apeptide moiety cleavable by one or more enzymes (e.g., matrixmetalloproteinase (MMP)).

In certain embodiments, the hydrophilic branched polymers are linked toNCX or CX via one or more stable linkages that are resistant tohydrolysis or enzyme cleavage. Exemplary stable linkages include amide,C—C(carbon-carbon single bond), C═C (carbon-carbon double bond), C≡C(carbon-carbon triple bond), ether (C—O—C), urethane (carbamate)linkages.

In certain embodiments, the hydrophilic polymers are branchedpolyethylene glycol (PEG), for example, with 3, 4, 5 or more arms orbranches.

Any suitable covalent crosslinking (CX) may be utilized. In certainembodiments, the covalent crosslinking (CX) is formed by copper-free,strain-promoted azide-alkyne cycloaddition or copper-catalyzedazide-alkyne cycloaddition. In certain embodiments, the covalentcrosslinking (CX) is formed by a click chemistry coupling between:

wherein R₁ is a group comprising —N₃, X is selected from ester andcarbonate groups or is absent, and each n is independently an integerfrom about 1 to about 400 (e.g., from about 10 to about 400, from about100 to about 400, from about 200 to about 400, from about 1 to about200, from about 1 to about 100, from about 1 to about 50, about 50 toabout 100, from about 1 to about 10, from about 3 to about 100, fromabout 3 to about 50, from about 3 to about 20, from about 3 to about10); and

wherein R₂ is

or a group comprising a cyclic or acylic alkyne group, Y is selectedfrom —NH— and —O— groups or absent, and each m is independently aninteger from about 1 to about 400 (e.g., from about 10 to about 400,from about 100 to about 400, from about 200 to about 400, from about 1to about 200, from about 1 to about 100, from about 1 to about 50, about50 to about 100, from about 1 to about 10, from about 3 to about 100,from about 3 to about 50, from about 3 to about 20, from about 3 toabout 10).

In certain embodiments, R₂ is

wherein R₃ is a group comprising a group comprising a cyclic or acyclicalkyne group, each of p and q is an integer from about 1 to about 6(e.g., 1, 2, 3, 4, 5, 6).

In certain embodiments, R₃ comprises a group selected fromdibenzylcyclooctyne (DBCO), dibenzocyclooctyne-amine,dibenzocyclooctyne-N-hydroxysuccinimidyl ester,(1R,8S,9s)-Bicyclo[6.1.0]non-4-yn-9-ylmethanol,(1R,8S,9s)-bicyclo[6.1.0]non-4-yn-9-ylmethyl N-succinimidyl carbonate,and dibenzocyclooctyne-maleimide groups.

In certain embodiments, the covalent crosslinking (CX) is formed by aclick chemistry coupling of:

and the non-covalent crosslinking (NCX) of hydrophobic andhydrogen-bonding interactions between un-coupled DBCO moieties:

The degree of covalent crosslinking (CX) vs non-covalent crosslinking(NCX) may be adjusted as needed. In certain embodiments, about 100% to10% (e.g., about 100% to 20%, about 100% to 30%, about 100% to 40%,about 100% to 50%, about 90% to 10%, about 80% to 20%, about 70% to 20%,about 70% to 30%, about 70% to 40%) are covalent SPAAC crosslinks and 0%to 90% (e.g., about 0% to about 80%, about 0% to about 70%, about 0% toabout 60%, about 0% to about 50%, about 10% to about 90%, about 20% toabout 80%, about 30% to about 80%, about 30% to about 70%, about 30% toabout 60%) are DBCO-DBCO non-covalent crosslinks.

The polymer content may also be adjusted as needed. In certainembodiments, the polymer content is from about 1% to about 20% (e.g.,from about 1% to about 15%, from about 1% to about 10%, from about 1% toabout 8%, from about 1% to about 5%, 2.5% to about 10%, from about 2.5%to about 8%, from about 2.5% to about 7%, from about 2.5% to about 6%,from about 2.5% to about 5%, from about 5% to about 10%, from about 10%to about 15%, from about 15% to about 20%).

The hydrogel of the invention may be prepared to exhibit pre-selectedproperties such as stiffness, plasticity/viscoelasticity and degradationcharacteristics. For example, the hydrogel may be characterized by oneor more of followings: a compressive stiffness from about 0.2 KPa toabout 20 KPa (e.g., from about 0.2 KPa to about 15 KPa, from about 0.2KPa to about 12 KPa, from about 0.2 KPa to about 10 KPa, from about 1KPa to about 20 KPa, from about 1 KPa to about 15 KPa, from about 1 KPato about 10 KPa, from about 3 KPa to about 20 KPa, from about 5 KPa toabout 20 KPa); a swelling ratio from about 1.5 to about 150 (e.g., fromabout 15 to about 120, from about 15 to about 100, from about 15 toabout 70, from about 15 to about 50, from about 25 to about 150, fromabout 50 to about 150, from about 70 to about 150); or a disintegrationrate from about <2 days to about >1 year (e.g., from about 2 days toabout 250 days, from about 2 days to about 150 days, from about 2 daysto about 60 days, from about 2 days to about 30 days, from about 3 daysto about 14 days, from about 7 days to about 1 year, from about 14 daysto about 1 year). The hydrogel of the invention may be prepared toexhibit viscoelasticity as characterized with any degrees ofcreep/stress-relaxation behavior.

In another aspect, the invention generally relates to a hydrogelcomposition comprising a hydrogel disclosed herein and a biologicallyactive payload encapsulated therein.

In certain embodiments, a partial or complete de-crosslinking of thehydrogel partially or completely releases the biologically activepayload.

Any suitable biologically active payload may be encapsulated.

In certain embodiments, the biologically active payload comprises cells.In certain embodiments, the cells are mammalian cells or cell aggregates(e.g., cell pellets and/or organoids) selected from embryonic stemcells, induced pluripotent stem cells, mesenchymal stem cells, bonemarrow stromal cells, hematopoietic stem cells, osteoblasts,chondrocytes, endothelial cells, epithelial cells, myoblasts, periostealcells, beta cells, neutral cells, any cells differentiated from variousembryonic and adult stem cells, or cell lines.

In certain embodiments, the biologically active payload comprises abiomolecule selected from the group consisting of proteins, growthfactors, cytokines, and chemokines. In certain embodiments, thebiologically active payload comprises the biologically active payloadcomprises a bone morphogenetic protein-2, 4, 6, 7 or 2/7 heterodimer(BMP-2, BMP-4, BMP-6, BMP-7, BMP-2/7).

In certain embodiments, the biologically active payload comprises thebiologically active payload comprises a transforming growth factor (TGF)beta, stromal cell-derived factor 1 (SDF1), Indian hedgehog homolog(Ihh), fibroblast growth factor (FGF), insulin-like growth factor (IGF),or vascular endothelial growth factor (VEGF), or various neuralinduction factors.

In certain embodiments, the biologically active payload comprises abiomolecule selected from the group consisting of a nucleic acid, genevector, bioactive lipid factor, such as sphingosine-1-phosphate (SIP),and bacterial phage.

In certain embodiments, the biologically active payload comprises amineral (e.g., calcium apatites, calcium phosphates, hydroxyapatite, andsubstituted hydroxyapatites).

The hydrogel composition of invention are designed to be suitable foruse in situ delivery of a variety of agents, for example, therapeutic ordiagnostic agents, tissue repair and regeneration implants.

The hydrogel composition of the invention may be prepared to exhibitpre-selected properties such as stiffness, plasticity and degradationcharacteristics. For example, the hydrogel composition may becharacterized by one or more of followings: a compressive stiffness fromabout 0.2 KPa to about 20 KPa (e.g., from about 0.2 KPa to about 15 KPa,from about 0.2 KPa to about 12 KPa, from about 0.2 KPa to about 10 KPa,from about 1 KPa to about 20 KPa, from about 1 KPa to about 15 KPa, fromabout 1 KPa to about 10 KPa, from about 3 KPa to about 20 KPa, fromabout 5 KPa to about 20 KPa); a swelling ratio from about 15 to about150 (e.g., from about 15 to about 120, from about 15 to about 100, fromabout 15 to about 70, from about 15 to about 50, from about 25 to about150, from about 50 to about 150, from about 70 to about 150); or adisintegration rate from about 2 days to about 1 year (e.g., from about2 days to about 250 days, from about 2 days to about 150 days, fromabout 2 days to about 60 days, from about 2 days to about 30 days, fromabout 3 days to about 14 days, from about 7 days to about 1 year, fromabout 14 days to about 1 year). The hydrogel of the invention may beprepared to exhibit a viscoelasticity as characterized with any degreeof creep/stress-relaxation behavior. The hydrogel or hydrogelcomposition of the invention are preferably cytologically compatible.

In yet another aspect, the invention generally relates to a device orimplant that includes a hydrogel or a hydrogel composition disclosedherein.

The device or implant may be any suitable medical device or implant, foreither human or animal use. Exemplary devices and implants include 3D invitro tissue models or synthetic niche for 3D cell expansion.

In yet another aspect, the invention generally relates to a method formodulating one or more properties of a hydrogel. The method includes:incorporating in the hydrogel a 3-dimentional crosslinking network of acombination of covalent crosslinking (CX) and non-covalent crosslinking(NCX) of hydrophilic branched polymers,

adjusting and/or controlling the ratio of covalent crosslinking (CX) tonon-covalent crosslinking (NCX) of hydrophilic branched polymers; andadjusting and/or controlling the placement and ratio of one or morelabile linkages to one or more stable linkages in the hydrophilicbranched polymers. The hydrophilic branched polymers are star-branchedwith at least 3 arms. The property is one or more selected fromviscoelasticity, stiffness and degradation.

In certain embodiments, the ratio of covalent crosslinking (CX) tonon-covalent crosslinking (NCX) is from about 1:10 to about 10:1 (e.g.,from about 1:8 to about 8:1, from about 1:5 to about 5:1, from about 1:3to about 3:1, from about 1:2 to about 2:1, about 1:1).

In certain embodiments, the ratio of one or more labile linkages to oneor more stable linkages is from about 1:0 to about 0:1 (e.g., from about1:0.2 to about 0.2:1, from about 1:0.5 to about 0.5:1, from about 1:0.2to about 0.2:1).

In yet another aspect, the invention generally relates to a method fordelivering a biologically active payload. The method includes: placingin a subject in need thereof a device or implant comprising a hydrogelor hydrogel composition disclosed herein, and causing a controlledrelease of the biologically active payload.

This present invention presents a new approach for precisely tuninghydrogel stiffness and matrix plasticity via the explicit control overthe degree of covalent SPAAC crosslinking vs. dynamic physicalcrosslinks between untethered end groups of biorthogonal macromerbuilding blocks. Conventional methods of modulating synthetic hydrogelstiffness by increasing polymer fractions or covalent crosslinkingdegrees could be detrimental to the proliferation and long-termviability of encapsulated cells. Modulation of viscoelasticity ofnatural polymers such as alginate-based hydrogels involve the use ofalginate of different molecular weights, adjustment of degrees ofCa²⁺-crosslinking, and the covalent attachment of other polymer tethers.In contrast, with only 2 pairs of designer building blocks (4-armPEGend-functionalized with DBCO or azide via stable or labile linkers), thepresent invention offers ClickGels with a broad range of stiffness,plasticity and degradative properties could be prepared by simplyaltering their mixing ratio along with cells of interest.

By introducing robust yet dynamic physical crosslinks between untetheredDBCO end groups at the expense of reduction in SPAAC crosslinks,enhancement in both ClickGel stiffness and network plasticity can beachieved, permitting more robust proliferation and matrix deposition ofencapsulated chondrocytes.

The invention additionally allows controlled degradation in ClickGels toachieve predictable disintegration over a broad range by strategicplacement of a single labile linkage on either side of the SPAACcrosslink, and that ClickGel degradation promoted both proliferation andchondrogenic matrix deposition of encapsulated mouse or humanchondrocytes. Finally, timed release of chondrocytes from degradableClickGels was accomplished without negatively affecting the viability orchondrogenic phenotypes of released chondrocytes.

These properties combined make ClickGels, a wholly synthetic platformfree of biological contaminants and with readily and reproduciblytunable physical and mechanical properties, uniquely suited as 3Dsynthetic niches for chondrocyte encapsulation, in vitro expansion andtimed release. It could also benefit the ex vivo expansion of scarcelyavailable stem cells (e.g., hematopoietic stem cells) known to bedifficult to expand/enrich via conventional 2D cultures for othercell-based therapies. More broadly speaking, the novel concept ofengineering network viscoelasticity via controlled integration ofdynamic physical crosslinks and the strategic placement of single labilelinkages near crosslinking sites provide exciting new tools forengineering 3D cellular niches and tissue models for regenerativemedicine and drug discovery applications.

The following examples are meant to be illustrative of the practice ofthe invention, and not limiting in any way.

Examples

Partially or perfectly ([DBCO]:[N₃]=1) SPAAC-crosslinked non-degradableClickGels (FIG. 1) with tunable compressive stiffness and swellingbehavior were prepared by systemically altering the molar ratios ofnon-degradable 4-armPEG-azide (Y=absent) to non-degradable4-armPEG-amide-DBCO (X═NH) mixed in PBS (5 w/v % polymer content), forexample, from 1:0.6 to 0.6:1.

The partially SPAAC-crosslinked ClickGels formed with an excess of4-armPEG-azide ([DBCO]:[N₃]<1) exhibited much higher swelling ratios(FIG. 2a ) and significantly weaker compressive moduli (FIG. 2b ) thanthe perfectly SPAAC-crosslinked ClickGel. The observed 2-fold increasein swelling ratio and 2-fold decrease in compressive stiffness (at the0-30% strain range) when [DBCO]:[azide] changed from 1:1 to 0.6:1 can beattributed to the reduced percentage of SPAAC covalent crosslinks, thusa more loosely tethered 3D ClickGel network.

By contrast, partially SPAAC-crosslinked ClickGels formed with an excessof 4-armPEG-DBCO macromers exhibited significantly higher compressivemoduli and much lower swelling ratios compared to the perfectlySPAAC-crosslinked ClickGel (FIGS. 2a & 2 b). The fold decrease inswelling ratio and 2-fold increase in compressive stiffness can beattributed to the increasing physical crosslinks between untetheredDBCO's when [DBCO]: [N₃] changed from 1:1 to 1:0.6.

These physical crosslinks, presumably driven by a combination ofhydrophobic interaction between the tricycles of DBCO's and H-bondinginteractions between adjacent amide linkages (FIG. 1), were robustenough to overcome the reduction in covalent SPAAC crosslinks, resultingin a mechanically strengthened 3D network. The mechanical contributionof DBCO-DBCO physical crosslinks within the imperfectlySPAAC-crosslinked system (e.g., [DBCO]:[N₃]=1:0.6) was validated by thereduction in compressive moduli upon addition of polyaromatic dye todisrupt the DBCO-DBCO interaction (day 1, FIG. 2c ). Furthermore, theability of untethered DBCO groups to reform physical crosslinks wasevidenced by the restoration of the compressive moduli of the hydrogelupon dye removal (after 7-day equilibration in PBS) to the level ofClickGels without dye treatment (FIG. 2c ), supporting thedynamic/reversible nature of the DBCO-DBCO crosslinks. By contrast, theaddition and removal of the polyaromatic dye to and from the imperfectlySPAAC-crosslinked ClickGels with azides in excess (e.g.[DBCO]:[N₃]=0.6:1) did not cause significant perturbations in theirstiffness (FIG. 2c ), supporting negligible physical crosslinks betweenuntethered azide-terminated chains. Further demonstrating the robustnessof DBCO-DBCO physical crosslinks was the observation that gellingoccurred at a mismatched ratio as drastic as [DBCO]: [N₃]=1:0.3, withthe resulting ClickGel possessing only 30% covalent SPAAC crosslinks but70% DBCO-DBCO physical crosslinks stiffer than those formed at [DBCO]:[N₃]=0.6:1 and 0.8:1 (FIG. 8). By contrast, formulations with asignificant fraction of excess azide endgroups ([DBCO]: [N₃]=0.3:1,0.4:1 or 0.5:1) could barely gel into network with sufficient integritydue to lack of physical crosslinking among excess azide groups, despitethe 30-50% of covalent SPAAC crosslinks.

Viscoelastic hydrogels are attractive for cell encapsulation due totheir ability to better accommodate cell spreading, migration,proliferation and matrix deposition through more effective/faster stressrelaxation. Here we show that engineering the reversible DBCO-DBCOphysical crosslinks into the wholly synthetic network translated intosignificantly faster stress relaxation in these stiffer hydrogels (FIGS.2d & 2 e). Specifically, stress relaxation (τ_(1/2)) in hydrogels formedwith excess DBCO-terminated macromers (e.g. [DBCO]:[N₃]=1:0.6, 131±52min) was significantly faster than hydrogels with 100%SPAAC-crosslinking (DBCO]: [N₃]=1:1, 291±85 min) or those formed withexcess azide-terminated macromers ([DBCO]: [N₃]=0.6:1, (320±88 min).These data support that dynamic DBCO-DBCO physical crosslinks (theirbreakage and reformation) are more effective in dissipating energy thanany potential physical interactions between the SPAAC crosslinks or theuntethered azide-terminated PEG arms in the weaker ClickGels. There islikely some level of hydrophobic interactions among the triazolemoieties of the SPAAC crosslinks, the disruption of which may haveexpedited energy dissipation to a degree comparable to that due to themobility of untethered PEG-azide arm (no statistically significantdifference in τ_(1/2), [DBCO]: [N₃]=1:1 vs. [DBCO]: [N₃]=0.6:1; FIG. 2e). It should be noted that the time scale of shear stress relaxationtime τ_(1/2) determined for these hydrogels were much faster, althoughtheir general trend remain the same. For instance, the shear stressrelaxation time τ_(1/2) of a ClickGel with [amide-DBCO]:[CH2-N₃]=1:0.6was determined as 1268.2±304.7 seconds (FIG. 2f ) as opposed to 131±52min under compressive mode on DMA.

Further validating the robustness of DBCO-DBCO physical crosslinks wasthe observation that gelling occurred at mismatched ratio as drastic as[DBCO]:[N₃]=1:0.3, which comprised only 30% covalent SPAAC crosslinksand 70% DBCO-DBCO physical crosslinks, resulting in a ClickGel stifferthan that formed at [DBCO]:[N₃]=0.6:1 or 0.8:1 (FIG. 8). By contrast,formulations with [DBCO]:[N₃]=0.3:1, 0.4:1 or 0.5:1 could barely gelinto network with sufficient integrity due to lack of physicalcrosslinking among excess azide groups, despite the 30-50% of covalentSPAAC crosslinks.

To determine a suitable cell encapsulation density within ClickGels, wemonitored the viability of iMACs encapsulated at an initial densityranging from 100000 to 1000000 cells per 25-μL perfectlySPAAC-cross-linked nondegradable ClickGel (5% w/v) and cultured inexpansion media over 8 weeks (FIG. 9a ). Whereas iMACs were able toproliferate within the ClickGel during the first 3-4 weeks at allencapsulation densities examined, the number of viable cells in thosewith very high initial encapsulation densities (750000 and 1000000 iMACsper gel) quickly declined after 6 weeks, with dead cells localizeddeeper within the gel due to possible nutrient deprivation. The initialencapsulation density of 250000 iMACs per gel resulted in the mostviable cells by 8 weeks (FIG. 9b ) while maintaining the expression ofchondrogenic markers type II collagen and aggrecan (FIG. 9c ). Thisencapsulation density was thus utilized to investigate the impact ofClickGel stiffness on resident cells.

Stiffer ClickGels with dynamic DBCO-DBCO physical crosslinks([DBCO]:[N₃]>1) supported better cell proliferation of the encapsulatediMACs over 4 weeks in EM compared to those encapsulated in the perfectlySPAAC-crosslinked ClickGel (FIG. 2a ), along with more robust type IIcollagen expression by iMACs in the stiffer gels (FIG. 2c ).

Using the optimized initial cell encapsulation density (25,000cells/25-μL gel) and polymer content of ClickGel (5 w/v %), we testedthe hypothesis that stiffer ClickGels with dynamic DBCO-DBCO physicalcrosslinks ([DBCO]:[N₃]>1) constitute a more adaptive/permissive nicheenvironment for cell proliferation and ECM deposition (FIG. 3a ).Indeed, the stiffer and more viscoelastic ClickGel strengthened byDBCO-DBCO crosslinks supported better cell proliferation of theencapsulated iMACs over 4 weeks in EM compared to those encapsulated inthe perfectly SPAAC-crosslinked ClickGel (FIG. 3b ), accompanied withmore robust type II collagen expression (FIG. 3d ). By contrast, in theabsence of DBCO-DBCO physical crosslinks, iMACs encapsulated within theweaker and less viscoelastic ClickGels ([DBCO]:[N₃]<1) exhibited poorercell proliferation and viability beyond the first week, with theClickGel with the most untethered azide-terminated PEG arms ([DBCO]:[N₃]=0.6:1) being the least favorable 3D environment for iMACs.Consistent with these observations, only the stiffer and moreviscoelastic iMAC-laden ClickGel ([DBCO]: [N₃]=1:0.6) exhibitedstatistically significant enhancement in stiffness, presumably due tomore robust ECM deposition, after 4-week culture in EM (FIG. 3c ).

As polymer content is also known to affect the mechanical properties ofcrosslinked hydrogels, we investigated whether and how modulatingClickGel polymer content along with the degree of SPAAC/physicalcross-links synergistically affect their compressive moduli and thecellular behavior of encapsulated iMACs. As expected, increasing anddecreasing polymer content by 2-fold (from 5 to 10 or 2.5% w/v) inperfectly SPAAC-cross-linked or stiffer ClickGels with excess DBCO'sresulted in proportional enhancement and reduction in compressivemoduli, respectively (FIG. 4a ). In the much weaker ClickGel formulationwhere azide residues were in large excess ([DBCO]:[N₃]=0.6:1),increasing or decreasing the polymer content failed to yieldstatistically significant changes in compressive moduli (<2-kPacompressive modulus). However, the benefit of increasing polymercontents in these weak ClickGels from 5 to 10% w/v was manifested bystatistically significant increase in early proliferation ofencapsulated iMACs within the first week and their subsequent viabilityover 4 weeks (by 1-4 fold increase; FIG. 4b ). In the perfectlySPAAC-cross-linked ClickGels, reducing polymer content from 5 to 2.5%w/v (compressive moduli from 3.9 to 2.5 kPa) negatively impacted boththe proliferation and long-term viability of encapsulated iMACs (FIG. 4c). Whereas iMACs encapsulated in 5% w/v perfectly SPAAC-cross-linkedClickGel were able to proliferate at least for 3 weeks, thoseencapsulated in 2.5% w/v ClickGel only proliferated within the firstweek and saw significant decline in viable cells starting week 2,further supporting that a minimal stiffness threshold for sustainedproliferation (3-4 weeks) is likely over 2.5 kPa. Interestingly, therewas no statistically significant benefit for the proliferation/viabilityof encapsulated iMACs by increasing polymer content from 5 to 10% w/v(compressive moduli from 3.9 to 7.5 kPa) in the perfectlySPAAC-cross-linked ClickGel (FIG. 4d ). This observation suggests thatonce the stiffness of a synthetic niche falls within a suitable range,further increasing the stiffness at the cost of increasing polymercontent (which may negatively impact nutrient/waste transport in and outof the 3D network) may not be beneficial. For the much stiffer yet moreviscoelastic ClickGel with substantial DBCO-DBCO physical cross-links([DBCO]:[N₃]=1:0.6), reducing polymer content from 5 to 2.5% w/v(compressive modulus from 7.9 to 3.3 kPa) slowed the proliferation after2 weeks and resulted in reduced overall viable cells by 4 weeks.Meanwhile, increasing its polymer content from 5 to 10% w/v (compressivemodulus from 7.9 to 11.8 KPa) did not benefit proliferation/viability ofencapsulated iMACs. The overall cell viability within the 10% w/vstiffest gel in fact reduced by 4 weeks, again supporting thatincreasing the stiffness beyond a certain range at the cost ofincreasing polymer content may not be beneficial. It is worth notingthat despite the similar compressive moduli of 10% w/v perfectlySPAAC-cross-linked (7.5 kPa) vs. 5% w/v DBCO-DBCO strengthened (7.9 kPa)ClickGels, the latter better supported sustained proliferation andviability of encapsulated iMACs over 4 weeks.

Increasing and decreasing polymer content of ClickGels resulted inproportional enhancement and reduction in their compressive moduli (FIG.4a ), but the benefit of increasing polymer contents on early cellproliferation (first week) was only manifested in the much weakerimperfectly SPAAC-crosslinked ClickGels with excess untethered azidechains ([DBCO]:[N₃]=0.6:1) (FIG. 4b ). In both perfectlySPAAC-crosslinked ClickGels and the stiffer and more viscoelasticClickGels with DBCO-DBCO crosslinking, reducing polymer content from 5to 2.5 w/v % (compressive moduli from 3.9 kPa to 2.5 kPa and from 7.9kPa to 3.3 kPa, respectively) negatively impacted both the proliferationand long-term viability of encapsulated iMACs (FIGS. 4c & 4 d). Therewas no significant benefit for increasing polymer content from 5 to 10w/v % in these ClickGels. This observation suggests that once thestiffness of a synthetic niche falls within a suitable range, furtherincreasing the stiffness at the cost of increasing polymer content(which could negatively impact nutrient/waste transport in and out ofthe 3D network) may not be beneficial. It is worth noting that despitethe similar compressive moduli of 10 w/v % perfectly SPAAC-crosslinked(7.5 kPa) vs. 5 w/v % partially SPAAC-crosslinked and DBCO-DBCOstrengthened more viscoelastic (7.9 kPa) ClickGels, the latter bettersupported sustained proliferation and viability of encapsulated iMACsover 4 weeks. Overall, these experiments reveal ˜3 kPa as a likely lowerthreshold for the ClickGel system below which the encapsulated iMACscould not undergo sustained proliferation. Beyond this threshold,introducing dynamic DBCO-DBCO physical crosslinks is advantageous topolymer content increases as a means to improve the stiffness of thechondrogenic cellular niche.

Interestingly, there was no statistically significant benefit for theproliferation/viability of encapsulated iMACs by increasing polymercontent from 5 to 10 w/v % (compressive moduli from 3.9 kPa to 7.5 kPa)in the perfectly SPAAC-crosslinked ClickGel. This observation suggeststhat once the stiffness of a synthetic niche falls within a suitablerange, further increasing the stiffness at the cost of increasingpolymer content (which may negatively impact nutrient/waste transport inand out of the 3D network) may not be beneficial.

For the much stiffer partially SPAAC-crosslinked ClickGel withsubstantial DBCO-DBCO physical crosslinks ([DBCO]:[N₃]=1:0.6), reducingpolymer content from 5 to 2.5 wt % (compressive modulus from 7.9 kPa to3.3 kPa) slowed the proliferation after 2 weeks and resulted in reducedoverall viable cells by 4 weeks. Meanwhile, increasing its polymercontent from 5 to 10 w/v % (compressive modulus from 7.9 kPa to 11.8KPa) did not benefit proliferation/viability of encapsulated iMACs. Theoverall cell viability within the 10 w/v % stiffest gel in fact reducedby 4 weeks, again supporting that increasing the stiffness beyond acertain range at the cost of increasing polymer content may not bebeneficial.

It was observed that despite the similar compressive moduli of 10 w/v %perfectly SPAAC-crosslinked (7.5 kPa) vs. 5 w/v % partiallySPAAC-crosslinked and DBCO-DBCO strengthened (7.9 kPa) ClickGels, thelatter better supported sustained proliferation and viability ofencapsulated iMACs over 4 weeks. It is likely that the reversible natureof the physical crosslinks between untethered DBCOs translated into amore dynamic/adaptable local environment to accommodate the expandingspace and help dissipate the local energy/strain imposed by theproliferating cell mass.

Overall, these experiments reveal that 3.3 kPa as a likely lowerthreshold compressive modulus for the ClickGel system below which theencapsulated iMACs would be unable to undergo sustained proliferationbeyond 2 weeks. ClickGels with compressive moduli of 3.9-11.8 kPa areable to support sustained proliferation and cell viability over 3-4weeks. Within this range, introducing dynamic DBCO-DBCO physicalcrosslinks is advantageous to increasing polymer content (which couldnegative impact nutrient/waste transport) as a means to improve thestiffness and introduce necessary matrix plasticity/viscoelasticity ofthe synthetic niche to enable sustained cell proliferation.

Using the 5 w/v % perfectly SPAAC-crosslinked and the stiffer and moreviscoelastic ClickGels containing various fractions of DBCO-DBCOphysical crosslinks ([DBCO]:[N₃]=1:1, 1:0.9, 1:0.8, 1:0.7 or 1:0.6), wethen validated the general applicability of the system for long-termencapsulation of human articular chondrocytes. The encapsulated hACs(500,000 cells per 25-pt gel) remained viable in all formulationsexamined over the 8-week culture in low-serum chondrogenic media (whichpromotes chondrogenic matrix synthesis rather than cell proliferation),supporting that nutrients/waste could readily penetrate in and out ofthese hydrophilic 3D network at this cell encapsulation density (FIG. 5a). Equally important, toluidine blue staining and immunofluorescentstaining revealed robust deposition of GAG and type II collagensecretions by hACs encapsulated in all ClickGel formulations examined,with minimal type X collagen detected by week 8 (FIG. 5b ; note that thehACs were isolated from the relatively healthy portion of discardedosteoarthritic joint tissues). The successful encapsulation of hACs byClickGels and their extended in vitro culture within these 3D syntheticniches without compromised viability or chondrogenic phenotype point topromising utilities for ex vivo cartilage tissue engineeringapplications.

For guided cartilage tissue regeneration in vivo, it is critical to alsodemonstrate the feasibility of tuning the degradative properties of theClickGel to promote the proliferation and chondrogenic matrix depositionof encapsulated cells and ensure their timely disintegration.Synchronizing the rate of synthetic niche degradation with that of theneo-tissue integration could preserve the overall mechanical integrityof the cell-laden construct throughout the dynamic guided tissueregeneration process. Too fast of a degradation will compromise thenecessary stiffness necessary for maintaining active metabolism ofencapsulated cells while too slow of a degradation will impede ECM (GAGand collagen fibrils are macromolecules of micron dimension) integrationand eventual replacement of the synthetic niche by regeneratedneotissue. (Nicodemus, et al. 2008 Tissue engineering. Part B, Reviews14 (2), 149; Chu, et al. 2017 Tissue Eng Part A 23 (15-16), 795.)

By mixing labile 4-armPEG-ester-azide and stable 4-armPEG-azidemacromers in varying ratios with the stable 4-armPEG-amide-DBCO whilekeeping [DBCO]:[total N₃]=0.7:1, 1:1 or 1:0.7, 5 w/v % ClickGels withvarying degrees of SPAAC and DBCO-DBCO physical crosslinking wereprepared. These ClickGels disintegrated in 18-53 days or remainedintact >150 days upon incubation in EM (FIG. 6a ). The weaker partiallySPAAC-crosslinked ClickGels (azide in excess) with the highest labile4-armPEG-ester-azide fraction (100%) disintegrated the fastest while themuch stiffer ClickGels with DBCO-DBCO physical crosslinks containing theleast fraction of labile 4-armPEG-ester-azide (60%) disintegrated theslowest. The hydrophobicity around the DBCO-DBCO physical crosslinkscombined with the more densely packed network has likely slowed freewater penetration to the labile ester linkage near the SPAAC crosslinks.At a given degree of SPAAC/physical crosslinking, the disintegrationrate expectedly accelerated with the increasing fractions of labile4-armPEG-ester-azide macromer. In contrast to tuning the molecularweight of degradable polymer chains, covalent crosslinking contents, oroverall polymer fractions as the means of altering degradable hydrogeldegradation rates, the prospectively tuning of ClickGel degradation rateby facile adjustment of the ratio of a pair of building blockscontaining a single labile linkage (without altering overall polymercontent) avoids excessive immunogenic acidic degradation products orpoor nutrient transport associated with high polymer fractions.

Next examined was the proliferation/viability of iMACs encapsulated inperfectly SPAAC-crosslinked 5 w/v % ClickGels with various fractions oflabile 4-armPEG-ester-azide over 4-week culture in EM. The iMACsencapsulated in faster-degrading ClickGels better proliferated andmaintained their viability throughout 4 weeks (FIG. 6b ), consistentwith the higher fraction of live cells in faster degrading ClickGels ata given degree of SPAAC-crosslinking as revealed by live/dead staining(FIG. 10). The faster-degrading ClickGels also supported more robusttype II collagen secretion by the encapsulated iMACs (FIG. 6c ). It isworth noting that the cell-laden construct with 100% labile4-armPEG-ester-azide did not fully disintegrate on day 28 (although someviable cells already release from the weakening gel). The slightlyslower disintegration compared to the cell-free construct (whichdisintegrated in 25 days in EM) is likely due to the high cell massimpeding free water penetration to some extent.

When the perfectly SPAAC-crosslinked 5 w/v % ClickGel with increasingfractions of labile 4-armPEG-ester-azide were used to encapsulate hACs,more robust GAG and type II collagen secretions by encapsulated hACswere observed in all degradable formulations after 4 weeks inchondrogenic culture (FIG. 5d ). The expression of hypertrophy markertype X collagen by hACs was not observed in the faster degradingformulations (75% and 100% 4-armPEG-ester-azide). These observationssupport that degrading synthetic niches are more conducive tochondrogenic matrix deposition by encapsulated chondrocytes in general.

Also examined was the tunable release of chondrocytes from degradableClickGels and whether the released cells maintain their chondrogenicphenotype, which are critical for matrix assisted autologous chondrocyteimplantation applications. Chondrocytes tend to lose their chondrogenicphenotype in monolayer cultures with increasing passages. Thus,chondrocyte proliferation within a degradable ClickGel niche andsubsequent release without compromising their chondrogenic phenotype(FIG. 7a ) could provide a promising solution.

Crystal violet staining was used to monitor the iMACs released (andadhered to the culture plate) from perfectly SPAAC-crosslinked ClickGelswith varying fractions of labile 4-armPEG-ester-azide over 30day-culture in EM (FIG. 7b ). On day 10, only the ClickGel formed with100% labile 4-armPEG-ester-azide released a small number of iMACs. Onday 15, this faster-degrading ClickGel released a bulk content ofencapsulated cells while the ones containing 75%, 50% or 25% labile4-armPEG-ester-azide began to release. On day 30, the fastest-degradingClickGel released the remaining iMACs whereas the one containing 75%labile ester-N3 linkages started to release the bulk content of itsencapsulated cells. These observations support a positive correlationbetween the content of labile ester linkages within the ClickGel (geldisintegration rate) and cell release rate.

When varying ratios of labile 4-armPEG-ester-DBCO (100%, 50%, 0%) andstable 4-armPEG-amide-DBCO were mixed with 100% labile 4-armPEG-ester-N₃to encapsulate iMACs, the same correlation was observed between labileester linkage fractions and cell release rate, although the cell-ladenconstructs fully disintegrated more rapidly, on days 25, 42, and 56,respectively.

The iMACs released from these ClickGels were then pelleted and culturedin EM for another 10 days before being stained for phenotypicalchondrogenic markers. Robust expression of type II collagen and GAG wasobserved with these iMAC pellets (FIG. 7c ), supporting that these cellsmaintained their chondrogenic phenotype throughout their encapsulationwithin the degradable ClickGels as well as after their release.

Experimental Macromer Synthesis

4-armPEG-amide-DBCO, 4-armPEG-ester-DBCO, 4-armPEG-azide and4-armPEG-ester-azide were synthesized from 4-armPEG20k, purified andcharacterized by ¹H and ¹³C NMR and GPC following published protocols.(Xu, et al. 2014 J. Am. Chem. Soc. 136 (11), 4105.)

ClickGel Preparation

All hydrogels were prepared by mixing azide-terminated andDBCO-terminated macromers in varying stoichiometric ratios in phosphatebuffered saline (PBS, pH 7.4). The resultant solution was transferredinto either a Teflon mold (10 mm diameter, 200 μL) for mechanical tests,onto a sterilized Parafilm for degradation test (25 μL) or swellingratio test (50 μL). The formulation was allowed to gel at roomtemperature up to 15 min (weaker formulations allowed to gel longer).For ClickGels consisting of 3 macromers, the labile and stableazide-terminated macromers or DBCO-terminated macromers were first mixedseparately.

Cell Encapsulation in ClickGels

A 25-4 suspension of cells, azide-terminated and DBCO-terminatedmacromers in varying stoichiometric ratios in PBS was mixed and pipettedonto sterilized Parafilm, and allowed to gel 2-30 min (weakerformulation allowed to gel longer) before being transferred intolow-attachment 24-well plates.

Equilibrated Swelling Ratio

As-prepared 5% w/v ClickGels (50 μL) were placed into 2 mL of 0.1-M PBS(pH 7.4) each and equilibrated on an orbital shaker for 1 week at roomtemperature. The fully equilibrated ClickGels were dabbed by KimWipe toremove excess aqueous buffer and weighed (W_(h)). The equilibratedweight swelling ratio was determined by the weight of the fully hydratedhydrogel (W_(h)) versus the weight of the dried specimen uponlyophilization (W_(d)) using the following equation:

weight swelling ratio=(W _(h) −W _(d))/W _(d)  Equilibrated

Compressive Moduli

Unconfined compressive test was performed on a dynamic mechanicalanalyzer (DMA800, TA Instruments) at 20° C. Cylindrical specimens (N=3)were compressed under the force controlled mode, ramping from 0.02 N to12 N at 5 N/min in PBS (pH 7.4) or DMEM (for cell-laden hydrogel) in asubmersion compression fixture. The slopes of the stress-versus-straincurves in the linear range of 0-30% or the higher 60-65% strain rangewere used for calculating the compressive moduli.

Stress Relaxation Time

The rates of stress relaxation, time taken to relax the stress to halfof initial loading (τ_(1/2)), of ClickGels were measured fromcompression tests of cylindrical specimens (N=7, pre-equilibrated in PBSfor 36 h) on a dynamic mechanical analyzer (DMA800, TA Instruments)equipped with a submersion compression fixture under stress relaxationmode. All specimens were applied with a 15% constant compressive strain(ramped within 0.1 min and held throughout the test) while the load wasrecorded over time.

Modulating Physical Cross-Links within ClickGels with Polyaromatic Dye

As-prepared 5% w/v ClickGels (200 μL, N=3) were placed into 2 mL of0.125-mg/mL Bromophenol Blue sodium salt solution (in PBS, pH 7.4, 1 M)and equilibrated on an orbital shaker at room temperature. After 24 h,the dye solution was removed and the ClickGels were imaged and subjectedto mechanical testing as described above. The ClickGels were thenequilibrated in 2-mL PBS, with daily replacement of fresh PBS for 7 daysbefore they were imaged and subjected to mechanical testing.

Immature Murine Articular Chondrocyte (iMAC) Isolation

iMAC was harvested from wild-type C57BL/6 neonates (4-6 days old) aspreviously described. (Huang, et al. 2016 PLoS One 11 (1), e0148088.)Briefly, the cartilage from knee and ankle joints were collected anddigested for ˜1 h in type II collagenase (Worthington) solution (3 mg/mLor 900 U/mL in high glucose DMEM). The cartilage pieces were then rinsedin PBS and transferred to a more dilute type II collagenase solution (1mg/mL or 300 U/mL) and incubated at 37° C. on a tube rotator for 5-6 h.The digested cartilage solution was then filtered through a 70-μm nylonmesh to obtain single cell suspension. The cells were pelleted andimmediately used for hydrogel encapsulation.

Human Chondrocyte (hAC) Isolation

Discarded tissues from osteoarthritic patient (age 66, male) undergoingtotal knee arthroplasty were collected and stored in ice cold PBSsupplemented with penicillin/streptomycin for less than an hour beforeprocessing. Using a sterilized razor blade, structurally intactarticular cartilage from the femoral condyles were shaved off and mincedinto tiny pieces. The cartilage pieces (3-5 g) were transferred to a50-mL conical tube with type II collagenase solution (lmg/mL or 300U/mL) and digested overnight at 37° C. on a rotator. Any residualcartilage fragments were removed by filtration through a 70-μm nylonmesh and the cell suspension was plated at a cell density of 25×10³/cm².Upon reaching 80% confluency, cells were trypsinized and immediatelyencapsulated in the hydrogels. Comparison of hAC matrix productionacross different hydrogel formulations were performed using cells from asingle donor.

Cell Culture

Cell-laden ClickGels were cultured at 37° C. with 5% CO₂ in eitherexpansion media (EM) or chondrogenic media (CM), with media changesevery 2-3 days. EM consisted of high glucose DMEM (Invitrogen)supplemented with 10% FBS (Gibco) and 1% penicillin/streptomycin(Corning). CM consisted of high glucose DMEM supplemented with 40 μg/mLL-proline (Sigma), 100 μg/mL sodium pyruvate (Sigma), 1%insulin-transferrin-selenous acid mixture (B&D Bioscience), 100 nMdexamethasone (Sigma) and 10 ng/mL TGF-β3 (R&D systems).

Monitoring Hydrogel Disintegration in EM

The ClickGel disintegration in EM was monitored at 37° C. in ahumidified incubator with 5% CO₂. As-prepared ClickGels (25 μL, 5% w/v)were placed in 750 μL of EM at 37° C., with change of fresh EM every 2-3days. The integrity of the ClickGel specimen was monitored daily. Thetime when the ClickGel completely disintegrated into the aqueous mediawas recorded as the disintegration time.

Cell Counting Kit-8 (CCK-8) Assay

The viability/proliferation of chondrocytes in each ClickGel formulation(n=3) was examined using CCK-8 assay (Dojindo), which is based on theconversion of a water-soluble tetrazolium salt, to a water-solubleformazan dye upon reduction by dehydrogenases in viable cells. (Han, etal. 2011 Antimicrob. Agents Chemother. 55 (10), 4519-4523.) At each timepoint, the cell-laden ClickGel was incubated with 500-μL 10% v/v CCK-8reagent in EM for 4 h before the media was transferred to a 96-wellplate (100 μL/well, n=3) for absorbance reading at 450 nm using aMultiskan FC microplate photometer (Thermo Scientific). Absorbance ofCCK-8 solution incubated in the absence of cell-laden ClickGel was readfor background subtraction.

Live/Dead Staining of Encapsulated Chondrocytes

Live (green)/dead (red) staining (Molecular Probes) of the cell-laden 3Dconstructs was performed with working solutions of 2-μM calcein AM and1-μM EthD-1 according to manufacturer's instructions. The stainedcell-laden ClickGels were mounted on a glass bottom dish and imagedusing a TCS SP5 II (Leica) confocal microscope. The composite image wascreated by overlaying 21 consecutive Z-stack images 5-μm apart.

Histochemical Staining for Glycosaminoglycan (GAG)

Cell-laden ClickGels retrieved from culture were fixed with 10% neutralbuffered formalin, serially dehydrated with ethanol and embedded inglycol methacrylate and sectioned into 5-um sections. Mounted sectionswere stained by toluidine blue to visualize the secreted sulfated GAG(in purple). At least 2 sections, 100 μm apart, were examined for eachconstruct.

Immunofluoresence Staining for Secreted Chondrogenic Matrix Proteins

Cell-laden ClickGels retrieved from culture were fixed by 10% neutralbuffered formalin and then washed thrice in in PBS with 1% BSA. Primaryantibody against type II collagen (Millipore, 1:200) or type X collagen(eBioscience, 1:200) was added for incubation at 37° C. for at least 1 hwith mild agitation. Negative controls were incubated in PBS withoutprimary antibodies. The ClickGels were washed thrice again with PBS (1%BSA). Alexa 488-conjugated goat secondary antibody against mouse(Molecular Probes, 1:200) was added for incubation at 37° C. for 1 hbefore washes in PBS (0.1% BSA, 3 times) in PBS and staining with DAPIreagent (Invitrogen) according to manufacturer's instructions. Thestained 3D hydrogels were first mounted on a glass bottom dish andimaged using a TCS SP5 II (Leica) confocal microscope. They were thenbisected to image the interior of the construct. Composite images werecreated by overlaying 21 Z-stack images 5-μm apart.

Cell Release Studies

iMACs were encapsulated in ClickGels formed between 4-armPEG-amide-DBCOand premixed 4-armPEG-ester-azide (0%, 25%, 50%, 75% and 100%) and4-armPEG-azide. Cell-laden ClickGels were cultured in EM. On days 10, 15and 30, cells released from the ClickGels and adhered to the cultureplates were fixed (10% formalin) and visualized by crystal violetstaining, while the cell-laden ClickGels were transferred to a newculture plate for continued culture.

In another subset of experiments, the cell-laden ClickGels were allowedto completely disintegrate within a 5 mL culture tube. The releasedcells were pelleted at 450×g for 10 min and cultured in EM for 10 days.The pellets were then fixed in 10% formalin, paraffin embedded,sectioned and stained with toluidine blue for GAG or for type IIcollagen (Millipore, 1:200) by immunohistochemistry withdiaminobenzidine (DAB) staining.

Statistical Analysis

One-way or two-way ANOVA with multiple comparisons test was performed asappropriate using Prism (GraphPad, version 7) to determine thestatistical significance between different ClickGel formulations (forswelling ratio, compressive modulus, and CCK-8 quantification at eachtime point), with p value less than 0.05 considered significant.

Applicant's disclosure is described herein in preferred embodiments withreference to the Figures, in which like numbers represent the same orsimilar elements. Reference throughout this specification to “oneembodiment,” “an embodiment,” or similar language means that aparticular feature, structure, or characteristic described in connectionwith the embodiment is included in at least one embodiment of thepresent invention. Thus, appearances of the phrases “in one embodiment,”“in an embodiment,” and similar language throughout this specificationmay, but do not necessarily, all refer to the same embodiment.

The described features, structures, or characteristics of Applicant'sdisclosure may be combined in any suitable manner in one or moreembodiments. In the description, herein, numerous specific details arerecited to provide a thorough understanding of embodiments of theinvention. One skilled in the relevant art will recognize, however, thatApplicant's composition and/or method may be practiced without one ormore of the specific details, or with other methods, components,materials, and so forth. In other instances, well-known structures,materials, or operations are not shown or described in detail to avoidobscuring aspects of the disclosure.

In this specification and the appended claims, the singular forms “a,”“an,” and “the” include plural reference, unless the context clearlydictates otherwise.

Unless defined otherwise, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art. Although any methods and materials similar or equivalent tothose described herein can also be used in the practice or testing ofthe present disclosure, the preferred methods and materials are nowdescribed. Methods recited herein may be carried out in any order thatis logically possible, in addition to a particular order disclosed.

INCORPORATION BY REFERENCE

References and citations to other documents, such as patents, patentapplications, patent publications, journals, books, papers, webcontents, have been made in this disclosure. All such documents arehereby incorporated herein by reference in their entirety for allpurposes. Any material, or portion thereof, that is said to beincorporated by reference herein, but which conflicts with existingdefinitions, statements, or other disclosure material explicitly setforth herein is only incorporated to the extent that no conflict arisesbetween that incorporated material and the present disclosure material.In the event of a conflict, the conflict is to be resolved in favor ofthe present disclosure as the preferred disclosure.

EQUIVALENTS

The representative examples are intended to help illustrate theinvention, and are not intended to, nor should they be construed to,limit the scope of the invention. Indeed, various modifications of theinvention and many further embodiments thereof, in addition to thoseshown and described herein, will become apparent to those skilled in theart from the full contents of this document, including the examples andthe references to the scientific and patent literature included herein.The examples contain important additional information, exemplificationand guidance that can be adapted to the practice of this invention inits various embodiments and equivalents thereof.

1. A hydrogel comprising a 3-dimentional crosslinking network of acombination of covalent crosslinking (CX) and non-covalent crosslinking(NCX) of hydrophilic branched polymers,

wherein the hydrophilic branched polymers are star-branched with atleast 3 arms.
 2. The hydrogel of claim 1, wherein the physicalinteractions are selected from the group consisting ofhydrophobic-hydrophobic interaction, π-π stacking, hydrogen bonding,electrostatic interaction, polar interaction, or a combination of one ormore thereof.
 3. The hydrogel of claim 1, wherein the hydrophilicbranched polymers are linked to NCX or CX via one or more labilelinkages and/or one or more stable linkages.
 4. The hydrogel of claim 3,wherein the one or more labile linkages are susceptible to hydrolysis.5. The hydrogel of claim 4, wherein the one or more labile linkagescomprise a group selected from the group consisting of ester, carbonate,orthoester, anhydride, and thioester.
 6. The hydrogel of claim 3,wherein the one or more labile linkages comprise a peptide cleavable byone or more enzymes.
 7. The hydrogel of claim 6, wherein the peptide isa matrix metalloproteinase (MMP) substrate cleavable by matrixmetalloproteinase (MMP).
 8. The hydrogel of claim 3, wherein the one ormore stable linkages are resistant to hydrolysis or enzyme cleavage. 9.The hydrogel of claim 8, wherein the one or more stable linkages areselected from the group consisting of amide, C—C, C═C, C≡C, ether,urethane linkages.
 10. The hydrogel of claim 1, wherein the hydrophilicpolymers are branched polyethylene glycol.
 11. The hydrogel of claim 10,wherein the branched polyethylene glycol comprises 4 arms.
 12. Thehydrogel of claim 1, wherein the covalent crosslinking (CX) is formed bycopper-free, strain-promoted azide-alkyne cycloaddition orcopper-catalyzed azide-alkyne cycloaddition.
 13. The hydrogel of claim12, wherein the covalent crosslinking (CX) is formed by a clickchemistry coupling between:

wherein R₁ is a group comprising —N₃, X is selected from ester orcarbonate groups or is absent, and each n is independently an integerfrom about 1 to about 400; and

wherein R₂ is

or a group comprising a cyclic or acylic alkyne group, Y is selectedfrom —NH— and —O— groups or absent, and each m is independently aninteger from about 1 to about
 400. 14. The hydrogel of claim 13, whereinR₂ is

wherein R₃ is a group comprising a group comprising a cyclic or acyclicalkyne group, each of p and q is an integer from about 1 to about 6.15-23. (canceled)
 24. The hydrogel of claim 1, characterized by one ormore of a compressive stiffness from about 0.2 KPa to about 20 KPa; aswelling ratio from about 15 to about 150; a disintegration rate fromabout 2 days to about 1 year.
 25. (canceled)
 26. A hydrogel compositioncomprising a hydrogel of claim 1 and a biologically active payloadencapsulated therein.
 27. The hydrogel composition of claim 26, whereina partial or complete de-crosslinking of the hydrogel partially orcompletely releases the biologically active payload. 28-42. (canceled)43. A device or implant comprising a hydrogel of claim
 1. 44-45.(canceled)
 46. A method for delivering a biologically active payload,comprising, placing in a subject in need thereof a device or implant ofany of claims 43-45; and causing a controlled release of thebiologically active payload. 47-57. (canceled)
 58. A method formodulating one or more properties of a hydrogel, comprisingincorporating in the hydrogel a 3-dimentional crosslinking network of acombination of covalent crosslinking (CX) and non-covalent crosslinking(NCX) of hydrophilic branched polymers,

adjusting and/or controlling the ratio of covalent crosslinking (CX) tonon-covalent crosslinking (NCX) of hydrophilic branched polymers; andadjusting and/or controlling the placement and ratio of one or morelabile linkages to one or more stable linkages in the hydrophilicbranched polymers, wherein the hydrophilic branched polymers arestar-branched with at least 3 arms; and the property is one or moreselected from viscoelasticity, stiffness and degradation. 59-60.(canceled)